Sutures pervade surgeries, but their performance is limited by the mechanical mismatch with tissues and the lack of advanced functionality. Existing modification strategies result in either deterioration of suture’s bulk properties or a weak coating susceptible to rupture or delamination. Inspired by tendon endotenon sheath, we report a versatile strategy to functionalize fiber-based devices such as sutures. This strategy seamlessly unites surgical sutures, tough gel sheath, and various functional materials. Robust modification is demonstrated with strong interfacial adhesion (>2000 J m−2). The surface stiffness, friction, and drag of the suture when interfacing with tissues can be markedly reduced, without compromising the tensile strength. Versatile functionalization of the suture for infection prevention, wound monitoring, drug delivery, and near-infrared imaging is then presented. This platform technology is applicable to other fiber-based devices and foreseen to affect broad technological areas ranging from wound management to smart textiles.
Sutures are a class of fiber-based devices primarily to mechanically approximate tissues or attach wearable/implantable devices to human body (1). They are in the form of either mono- or multifilaments (i.e., braided) and designed to degrade or stay permanently in the body. A variety of materials have been invented and adopted as surgical sutures, including plastics (degradable: polyglycolide and polylactic acid; nondegradable: nylon and polypropylene), biologically derived proteins (collagen and silk), and metals (stainless steel and nitinol). The sutures have been widely used in many branches of medicine such as wound closure and anastomosis with a global market of over 5 billion U.S. dollars (2). Despite the recent progress of tissue adhesives (3–7), they will remain indispensable for general surgical procedures because of their reliable performance, ease of implementation, and the capacity to exert larger forces than any tissue adhesives (2).
However, the performance of existing sutures has been limited by their poor biomechanical properties and lack of functionality, which are implicated in surgical and postsurgical complications. First, sutures are made of rigid dry materials (elastic modulus of >1 GPa) in contrast to soft hydrated tissues (elastic modulus of <100 kPa), as they need to carry substantial mechanical loading along the axis to approximate tissues (8). This mechanical mismatch is found to cause inflammation and impaired healing outcomes (9). Second, the rough surface of sutures, particularly for the braided sutures, can drag and rub against the contacting tissue during and after suture placement. This mechanical irritation can damage fragile tissues and those under disease conditions such as aneurysm and ulcer, leading to tissue dissection and other postsurgery complications (10, 11). In addition, clinically used sutures lack advanced functionality for wound management. Thus, multifunctional sutures are in demand to perceive, report, and respond to the wound healing process, for instance, delivering therapeutic to promote wound healing (12) and preventing surgical site infections (13). These functional sutures are developed recently, which feature drug delivery or sensing capacities. There are also drug-eluting or antibacterial sutures commercially available (e.g., Coated VICRYL Plus Antibacterial, Ethicon), capable of releasing drug or antibacterial compounds (14). However, limitations to these approaches remain, including complex fabrication process, high cost, limited physical integrity, and the abovementioned biomechanical constraints. These issues associated with surgical sutures are also found in other fiber-based devices, particularly those interfacing with the human body such as guidewires and smart textiles. New strategies to improve the biomechanical properties, functionality of sutures, and other fiber-based devices continue to be sought.
General strategies to functionalize sutures include bulk modification and surface functionalization. The former involves bottom-up approaches to (re)produce the suture (e.g., electrospinning and melt extrusion), which may compromise the suture’s strength and are inapplicable to commercially available sutures (15, 16). To minimize the alteration of the bulk properties, the surface functionalization is appealing, which results in a suture coating via dip coating/soaking (17, 18), layer-by-layer deposition (19, 20), grafting (21–23), and impregnation (24, 25). However, the suture coating is often weak and vulnerable to fragmentation and delamination, due to the chemical inertness of suture materials and the demanding mechanical loading of the suture application (e.g., shear and compression during suturing and knotting). The mechanical failure of suture coating results in the loss of functionality (fig. S1, A to C) and other side effects (e.g., burst drug release for drug-eluting suture coating). Evidently, the toughness and adhesion of the suture coating is thus mission critical and recognized as a prerequisite of any reliable functionalization.
For surface functionalization of surgical sutures, hydrogel technologies are promising in light of recent developments of tough hydrogel adhesion on various materials such as tissues (3), hydrogels (26), metals (27), and elastomers (28). These established strategies exploit preformed hydrogel patches to be applied on flat surfaces and thus are not compatible with fiber-based devices as sutures. Formation of hydrogel coating on complex structures has been recently reported but is focused on single-network hydrogels and elastomer-based devices (29). Neither this strategy nor the other methods are applicable on surgical sutures because of the chemical innerness of suture materials. In addition, most adhesive hydrogels are inclusion free and lack advanced functionality. Particularly, few works to date demonstrate wound bed monitoring and near-infrared (NIR) bioimaging applications with hydrogel coatings. Further development is needed to reinvent and repurpose the hydrogel coating for sutures and other fiber-based devices.
To address the abovementioned issues, we report a bioinspired design and fabrication method for multifunctional tough gel–sheathed (TGS) sutures. Different from the previously reported methods based on surface absorption or single-network hydrogels, our strategy features a double-network TGS strongly bonded with surgical sutures for robust modification. The TGS is hypothesized to help mitigate the mechanical mismatch and irritation of sutures when interfacing with tissues and to further provide a robust and versatile platform to functionalize commercially available sutures for advanced functionality. As a proof of principle, motivated by the clinical need of wound management, we will demonstrate the TGS sutures loaded with an antibacterial compound, pH-sensing microparticles, drugs, and fluorescent nanoparticles (NPs) for anti-infection, wound bed monitoring, drug delivery, and NIR bioimaging applications. This work will demonstrate that the TGS suture could unite the merits of suture fibers, tough hydrogels, and functional materials by design and thus achieve a unique combination of enhanced biomechanical properties and multiple functionality, which would be beneficial for general surgical procedures and wound management (fig. S1D).
Design and fabrication of TGS sutures
The design of TGS suture is inspired by endotenon sheath of tendon, which encapsulates and glues collagen fibers together (Fig. 1A) (30). The endotenon sheath is mechanically tough and strongly adhesive on the collagen fibers, attributing to its double-network structure: The hyaluronan-proteoglycan network binds with the collagen fibers, while the elastin network strengthens and toughens the whole structure. The endotenon sheath not only forms a frictionless surface of the tendon but also comprises cells and blood vessels for mass transport and repair of the tendon. Learning from the structure and function of the endotenon sheath, we propose a TGS to modify and further functionalize commercially available sutures. To achieve the strong bonding with the suture, the TGS contains one anchoring network that forms covalent bonds on the suture surface for strong adhesion and intertwines with another toughening network consisting of physical bonds, which can effectively dissipate energy for high toughness (general design strategy illustrated in Fig. 1B). This strategy is distinct from the surface absorption, the single-layer polymer grafting, and the hydrogel skin method that requires swelling and doping the substrate with free-radical initiators to form a single-network hydrogel coating (20, 23, 29).
Schematics of the structural and material design of (A) tendon and (B) TGS suture. Scanning electron microscope images of (C) TGS suture and (D) a zoom-in at the suture-sheath interface. Scale bars, 100 (C) and 25 μm (D). (E) Bright- field image of TGS suture. Scale bar, 500 μm. (F) A continuous stitch applied on porcine skin using TGS suture. Scale bar, 1 cm. Photo credit: Zhenwei Ma, McGill University.
The TGS sutures are prepared directly from commercially available surgical sutures with a facile two-step method. We illustrate the essential procedure with a widely used braided surgical suture, Coated VICRYL (Ethicon), consisting of polyglactin 910 [copolymer of 90% glycolide and 10% lactide (PLGA)] coated with polyglactin 370 and calcium stearate (14), and an alginate-polyacrylamide (PAAm) hydrogel that acts as the TGS because of its high toughness and excellent biocompatibility (31). As the pristine suture lacks the functional groups for anchoring, the suture was first treated with 1 M NaOH solution to create carboxylic acid groups on the surface, which was later primed with primary amine–rich chitosan macromolecules and coupling reagents [1-ethyl-3-(-3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS)] to form the anchoring network. To form the toughening network, we then inserted the modified suture into a glass capillary tube filled with a precursor solution of alginate-PAAm hydrogels; after gelation overnight, the TGS suture was obtained after post–cross-linking in 0.1 M CaCl2 solution (fig. S2). The applicability of this strategy to other hydrogels and suture materials will be presented below. Sutures of a wide range of diameters (0.01 to 1 mm) are used by surgeons in different surgical procedures. Our results demonstrated that the suture was robustly integrated with a thin TGS and that the thickness was tunable with the diameter of the capillary tube (Fig. 1, C to E). With our current setup, we can reproducibly fabricate 10- to 15-cm-long TGS suture threads with consistent coating thickness (fig. S3A). For the long-term storage of TGS sutures, they can be freeze-dried, kept at their dry state as other commercialized sutures, and then simply rehydrated in saline solutions before usage (fig. S3B). As a demonstration of its physical integrity, a continuous stitch with secured knotting was successfully performed on porcine skin using the TGS suture (Fig. 1F). The in vitro biocompatibility of the TGS suture–conditioned medium is comparable to that of the control medium, showing no notable difference in the in vitro viability of human vocal fold fibroblasts after 48-hour culture (fig. S4).
Strong suture-sheath bonding
To interrogate the bonding between the suture and the hydrogel sheath, we invented a pull-out test to characterize the adhesion energy of suture-sheath interface. Briefly, the suture was embedded within a hydrogel cuboid and then pulled out with an Instron machine (model 5965), while the force F and the displacement δ were recorded (fig. S5). The two opposing sides of the hydrogel cuboid were glued to two acrylic sheets as rigid constraints. It was observed that the adhesion survived at a large pull-out displacement (20 mm versus 30 mm of the adhesion interface) and that part of the tough gel matrix still attached to the suture even when it was completely pulled out (Fig. 2A and fig. S5).
Representative force-displacement curves (A) and the adhesion energy (B) measured from the pull-out tests [inset in (A)] of gel-sheathed sutures formed with polyglactin 910 suture and different hydrogels. Alg, alginate; Chi, chitosan. (C) Adhesion energy as a function of NaOH treatment time. (D) FEM results of the normalized adhesion energy (Γ/Γ0) as a function of the ratio of the sheath thickness and the suture radius (rg/rs). The force-displacement curves of pull-out test (E) and the adhesion energy (F) of TGS sutures encompassing various suture materials, including polyglactin 910 (PLGA), plain gut, and nylon. Data reported as means ± SD for n = 3 independent experiments.
An analytical model was developed and applied to calculate the adhesion energy. The strain energy density stored in the hydrogel sheath Ug can be calculated from the force-displacement curve until the point when the interface failed via
, where rs denotes the suture radius; rtot denotes the summation of the hydrogel sheath thickness rg and suture radius, i.e., rtot = rg + rs; L denotes the jointed length. The critical energy release rate, i.e., adhesion energy, was calculated from Ug with the equation (see Materials and Methods for derivation and other details)
With the above equations, we calculated that the adhesion energy obtained from the alginate-PAAm–sheathed polyglactin 910 suture was over 1000 J m−2 (Fig. 2B), which was comparable with the tough adhesion of hydrogels achieved on tissues, elastomers, and metals.
To reveal the mechanism for the strong adhesion, we next separately delineated the effects of the anchoring and toughening networks. First, to evaluate the importance of the toughening network, the suture sheath was formed with brittle single-network hydrogels (e.g., alginate or PAAm hydrogels), which largely resembled the strategies reported previously (17, 29). Without a tough matrix, the suture-sheath interface is vulnerable to rupture (adhesion energy of <50 J m−2) (Fig. 2B). Second, to confirm the role of the anchoring network, we fabricated the sheathed sutures without NaOH treatment or surface priming (chitosan/EDC/NHS). Evidently, the lack of strong anchoring network led to very low adhesion energy (<50 J m−2) (fig. S6A). The result also implied that the mechanical interlocking of the hydrogel within the braided suture played a small role. The importance of amide-based interfacial bonds was further confirmed with a positive correlation between the duration of NaOH treatment and the adhesion energy (Fig. 2C). Over 2000 J m−2 adhesion energy was achieved with 10-min surface activation. It can be concluded that the strong adhesion is attributed to the synergy of the energy dissipation of the hydrogel matrix and the covalent bonding at the suture-sheath interface. It should be noted that strong adhesion was achieved with the standard coated polyglactin 910 sutures. Given the same chemistry of the polyglactin 370 coating and the bulk of the suture filament, similar results are expected for the uncoated versions.
Finite element modeling of suture-sheath adhesion
To further model the interfacial failure mechanism and aid the design of suture sheath, we developed a finite element model (FEM) to simulate the pull-out process of a suture from a hydrogel sheath. The toughening network was characterized with the combined Ogden and modified Ogden-Roxburgh model that accounts for the Mullins effect, while the anchoring network was modeled as cohesive elements with low (24 J m−2) and high (300 J m−2) intrinsic toughness values Γ0 (32). Our simulation showed that the adhesion energy Γ increased with the intrinsic toughness Γ0 (fig. S6, B and C), consistent with the observation of stronger adhesion after longer NaOH treatment (Fig. 2C). To study the effect of sheath thickness, we varied the hydrogel thickness in terms of normalized hydrogel radius rg/rs while keeping the intrinsic toughness Γ0 = 300 J m−2. As the adhesion energy Γ is shown to scale linearly with Γ0 (33), we normalized the adhesion energy Γ by the intrinsic toughness Γ0. Large normalized adhesion energy was observed (Γ/Γ0 > 2) in all tested conditions (Fig. 2D), indicative of a potent toughening effect of the TGS (31). We found a nonmonotonic correlation between Γ/Γ0 and rg/rs. After examining the shear stress distribution along the joint interface, we interpret the observation as follows (fig. S7, A to E). Given a thin sheath (rg/rs ≤ 2), the suture-sheath interface debonds simultaneously, and the volume of energy dissipating materials increases with the thickness of the suture sheath, leading to higher adhesion energy; when the sheath is even thicker, the energy dissipation is confined around the crack tip, and thus, the total amount of dissipated energy is reduced, as well as the adhesion energy.
Wide applicability of TGS design
The design and method for the TGS sutures are applicable to a variety of surgical sutures and hydrogels. Besides the synthetic degradable PLGA suture, we have successfully fabricated the TGS sutures using naturally derived degradable plain gut sutures and synthetic nondegradable nylon sutures (Fig. 2, E and F) with the same method. Furthermore, the suture sheath can be formed with various hydrogels as the toughening network, which can interpenetrate with the chitosan-based anchoring network. As an example, we formed a TGS composed of a chitosan-PAAm hydrogel. The efficacy of our method is evidenced with high adhesion energy of the sheath-suture interfaces in all the tested conditions. This study leads to a family of TGS sutures of varying chemical compositions and properties and demonstrates the versatility of the proposed design and method.
Enhanced biomechanical properties
We next demonstrated that the intrinsic biomechanical properties of TGS sutures could help mitigate the limitations of clinically used suture materials. The core-sheath structure could provide low-stiffness hydrogel surface to resolve the mechanical mismatch between the suture and local tissues, without sacrificing the tensile strength of the pristine sutures for wound closure. With carefully controlled suture hydrolysis via NaOH treatment, the obtained TGS suture retained high tensile strength (3 GPa), comparable to pristine sutures (Fig. 3A and fig. S8). However, prolonged surface treatment will lead to compromised suture strength, despite the enhanced interfacial adhesion (Fig. 2C and fig. S8). The surface elastic modulus of the TGS suture is around 7 kPa measured by atomic force microscopy equipped with a cell-sized spherical probe (Fig. 3B), compared to ultrahigh stiffness of ~68 MPa of the pristine suture (fig. S9). The activation treatment is strictly limited to the very superficial layer of sutures and yet provides sufficient functional groups for binding with the sheath. The mechanical mismatch between the suture and soft tissues is thus remedied by the soft hydrogel sheath, which may also mitigate the local stress concentration and enable a mechanical microenvironment favorable for tissue regeneration.
(A) Stress-strain curves of the pristine and TGS sutures (polyglactin 910). (B) Representative microindentation force-indentation depth curve measured on the TGS suture surface. AFM, atomic force microscope. (C) Schematic of the tissue drag test. (D) Representative drag force-displacement curves of the pristine and TGS sutures. (E) Drag coefficients of suture (pristine or TGS) interfacing with various tissues (heart, skin, and liver). (F) Schematic of the ex vivo friction test of the suture placed on articular cartilage, where the PDMS is used as a substrate. Representative friction force-sliding distance curves (G) and the calculated friction coefficients (H) of intact cartilage and sutured cartilage with the pristine or TGS suture. Data reported as means ± SD for n = 3 independent experiments; ***P < 0.001, by two-tailed, one-way analysis of variance (ANOVA) with Holm-Sidak post hoc comparison.
In addition to the low stiffness, TGS suture also provides a slippery surface when interfacing with the tissues, which could substantially reduce the tissue drag and friction that have been linked with microtrauma and tissue damage (34). When passing through the tissue, traditional sutures can drag and damage the tissue; after the placement, the rough surface of sutures, particularly braided ones, can cause constant friction and wear on the contacting tissue. This is particularly severe for tissues under constant friction and impact, such as articular cartilage, where surgical suturing has been associated with higher risk of osteoarthritis (10). To characterize the tissue drag, we performed a customized ex vivo drag test, mimicking the suturing process, to determine the drag coefficient of the TGS and pristine sutures on soft tissues such as skin, heart, and liver (Fig. 3, C to E). The results show that the drag coefficients of the TGS sutures are two to three times lower than that of pristine sutures (Fig. 3E). A characteristic stick-and-slip phenomenon was observed in the drag force-displacement profile of the pristine sutures but not in the case of the TGS sutures (Fig. 3D and fig. S10). We also developed a customized setup to characterize the friction induced by the sutures on porcine articular cartilage (Fig. 3F). Using polydimethylsiloxane (PDMS) as an artificial tissue substrate, we characterized the friction coefficient between PDMS and cartilage or sutured cartilage using the pristine or TGS sutures. Our results show that friction coefficient for TGS sutures was comparable to articular cartilage and significantly lower than the pristine sutures (Fig. 3, G and H).
Besides the excellent biomechanical properties inherent to the TGS, the functionality of the TGS sutures will next be engineered through inclusion of various functional materials. It is feasible as the TGS can serve as a versatile platform, warranted by the robust modification, to readily encapsulate and deliver small payloads along with the suture to the strategic site adjacent to a wound. As a proof of principle, we will load the suture sheath with an antibacterial compound, pH-sensing microparticles, a model drug, and fluorescent NPs and then demonstrate the effectiveness of the functionalization below.
The interfilament spacing within braided sutures may attract bacteria through capillary forces and host bacteria growth. Owing to the cell-repellent nature of PAAm, the TGS suture exhibited excellent antifouling property with significantly lower bacteria adhesion, for both Gram-positive (Staphylococcus aureus) and Gram-negative (Pseudomonas aeruginosa) bacteria that are closely associated with surgical site infection (Fig. 4, A and B) (35). By further loading the TGS with an antibacterial compound [benzalkonium chloride (BZK)] widely used in many consumer products, over 99% of the adhered bacteria were killed (Fig. 4C and fig. S11), thanks to the positively charged quaternary ammonium compound incorporated into the hydrogel matrix (35). This study demonstrates the antifouling and antimicrobial functions of the TGS suture and its potential for mitigating surgical site infection.
(A) Representative fluorescence images of live (green)/dead (red) assay of bacteria (P. aeruginosa and S. aureus) seeded onto the pristine or TGS suture. Scale bar, 10 μm. (B) Total number of bacteria adhesion on pristine or TGS sutures. (C) Over 99% bacteria were killed on TGS sutures loaded with BZK. Representative images (D) and quantitative color change assay (reflected in gray scale) (E) of pH-sensing TGS suture immersed in solution with various pH levels. Photo credit: Zhenwei Ma, McGill University. (F) Seven-day normalized cumulative release profile of FITC-BSA from the pristine or TGS suture. (G) BSA loading capacity of pristine or TGS suture. Data reported as means ± SD for n = 3 independent experiments; ***P < 0.001, by two-tailed, one-way ANOVA with Holm-Sidak post hoc comparison.
The suture sheath can be further functionalized to monitor the physiological signals during wound healing processes because of the strategic location of surgical sutures at the wound site. As a proof of concept, we loaded pH-sensing beads into the TGS as pH-monitoring sutures. The wound bed pH is a key indicator of the wound healing process: Healthy skin is often acidic, and for wound bed exposed to body fluid, the pH is usually around 7, while the pH of chronic wounds or infected wounds could go up to 10 (36, 37). Our results show that the pH-sensing sutures readily transduce the pH signal into color change, visible to the naked eye (Fig. 4, D and E). The monitoring could be continuous and essentially minimally invasive as the sutures naturally penetrate through the wounded tissue; the semipermeable hydrogel sheath allows for mass exchange at the molecular level (yet does not allow bacterial invasion). The new suture function would particularly benefit the monitoring of chronic wounds by nonprofessionals and in-time intervention when the pH level is abnormal. By integrating the antifouling and antibacterial functions, the new suture technology could be potentially used as point-of-care systems for the management of chronic wounds.
The TGS suture can also serve as a depot to locally deliver drugs to the wound site. The TGS can encapsulate the drug and regulate its release through the drug-matrix interactions. As a proof of concept, we loaded the TGS with fluorescein isothiocyanate (FITC)–bovine serum albumin (BSA) as a widely used model drug to form drug-eluting TGS sutures. We also dip-coated the pristine suture with the model drug as a control for comparison. A 1-week cumulative release profile shows that TGS suture presents higher encapsulation efficiency and longer release period, compared to the control suture using dip-coating strategy (Fig. 4, F and G). Together with the abovementioned demonstrations, we show that the TGS platform enables the diagnostic, monitoring, and therapeutic functions to address the clinical needs for wound management.
The intricate surgical procedures and delicate soft tissues call for the precise targeting and visualization of sutures during and after operation. NIR fluorescence imaging is an emerging biomedical imaging modality for use in both fundamental scientific research and clinical practice (38). By incorporating our recently developed NIR fluorescent NPs (39) (with fluorescence emission wavelength at 1250 nm in NIR-II: 1000 to 1350 nm; Fig. 5, A and B, and fig. S12) into the TGS, we further expand the potential of our platform for deep-tissue bioimaging via a customized ex vivo setup (Fig. 5C). The fluorescent NP–loaded TGS suture exhibits bright linear fluorescence along suture with dark background captured by an NIR camera under excitation of an 806-nm laser in NIR-I (700 to 950 nm), and the strong signal could still be observed when the suture was covered by porcine tissue with up to 3-mm thickness (Fig. 5D). The intense NIR fluorescence and high contrast demonstrate the deep-tissue penetration capability of the fluorescent suture. To our knowledge, this first demonstration of fluorescent sutures with high optical transparency in the NIR biological window would facilitate minimally invasive surgeries to localize the suture during implementation and later removal and benefit fluorescence image-guided surgery and postsurgery bioimaging/diagnostics (40).